1. Field of the Invention
The present invention concerns a method to control a magnetic resonance system to generate magnetic resonance exposures of an examination subject, as well as a control device for a magnetic resonance system and a magnetic resonance system with which such a method can be implemented.
2. Description of the Prior Art
In order to obtain magnetic resonance exposures—i.e. image data—from a region of the inside of a body of an examination subject, the body or the body part that is to be examined, must initially be exposed to an optimally homogeneous, static basic magnetic field (most often designated as a B0 field). The nuclear spins in the body are thereby aligned parallel to the direction of the B0 field (typically designated as the z-direction). Moreover, radio-frequency pulses (also called “magnetic resonance radio-frequency pulses” in the following) are radiated into the examination subject with radio-frequency antennas. The frequency of the radio-frequency pulses is in the range of the resonance frequency (known as the Larmor frequency) of the nuclei to be excited (normally hydrogen nuclei) in the present magnetic field. The magnetic flux density of these radio-frequency pulses is typically designated with B1. By means of these radio-frequency pulses, the nuclear spins of the atoms in the examination subject are excited such that they are deflected by what is known as an “excitation flip angle” (also generally abbreviated to “flip angle” in the following) out of their steady state, parallel to the basic magnetic field B0. The relationship between the field strength B1 and the flip angle α that is achieved with this is provided by the equation
                              α          =                                    ∫                              t                =                0                            T                        ⁢                          γ              ·                                                B                  1                                ⁡                                  (                  t                  )                                            ·                                                          ⁢                              ⅆ                t                                                    ,                            (        1        )            wherein γ is the gyromagnetic ratio which—for most magnetic resonance examinations—can be considered as a fixed material constant, and T is the active duration of the radio-frequency pulse. The nuclear spins then initially precess around the z-direction and relax again bit by bit. The rotation of the nuclear spins around the precession cone can be viewed as a macroscopic nuclear magnetization in the x/y-plane (perpendicular to the z-direction). The magnetic resonance signals generated in the relaxation are acquired as raw data by means of radio-frequency reception antennas, and ultimately the magnetic resonance images are reconstructed on the basis of the acquired raw data. The spatial coding takes place with the use of rapidly switched gradient magnetic fields that are superimposed on the basic magnetic field during the emission of the magnetic resonance radio-frequency pulses and/or the acquisition of the raw data.
For the relaxation, the excited nuclei require a characteristic decay time that is dependent on the chemical bond and the molecular environment in which the excited nucleus is found. The different tissue types therefore characteristically differ in their signal, which leads to varying signal strengths (brightnesses) in the resulting image. Differentiation is made between two different characteristic relaxation times, the longitudinal relaxation time T1 and the transverse relaxation time T2. The longitudinal relaxation time T1 is determined by interaction with the surrounding atoms in the lattice. The transverse magnetization MT decays due to what is known as spin-spin interaction with adjacent atoms with continuing time t after the end of the magnetic resonance radio-frequency pulse, according to the equation:
                              MT          ⁡                      (            t            )                          =                              MT            ⁡                          (              0              )                                ·                      ⅇ                          t                              T                2                                                                        (        2        )            
This temporal response is schematically presented in FIG. 1 (in arbitrary units=a.U.) for three substances with differing transversal relaxation times or, respectively, T2 time constants. The solid curve corresponds to a substance with long T2; the dashed curve corresponds to a substance with medium T2; and the dotted curve corresponds to a substance with very short T2. The T2 time constant (also abbreviated only as “T2” in the following) is very different for different substances. While this time constant is relatively long for some tissue types and fluids (up to 100 ms or more), for bones, teeth or ice it is only between 30 μs and 80 μs, for example.
In the case of specific diagnostic questions (for example in orthopedic applications), images that show only substances with short T2 are required for monitoring of cryo-ablations or for MR-PET or PET attenuation correction. To create an image in which only substances with short T2 are visible, substances with long T2 could be saturated via very long (˜50 ms) pre-pulses. Alternatively, a second echo with longer echo time and a difference image made of first and second echo could be generated.
Furthermore, sequences with very short echo times—known as UEZ sequences (UEZ=ultra-short echo times with TE<0.5 ms)—enable the presentation of substances with short T2 that are not visible with conventional sequences. Examples of UEZ sequences are UTE (Ultra Short Echo Time), PETRA (Pointwise Encoding Time Reduction with Radial Acquisition) or z-TE (Zero Echo Time). In these sequences, for the most part a hard δ-pulse is applied and after this a data acquisition is begun in free induction decay (FID). However, since given very short times the signals of all substances decay approximately equally (see echo time TE1 in FIG. 1), no T2 contrast can be generated with these. Only with longer times (see echo time TE2 in FIG. 1) does a larger difference exist between the individual substances, and the image is T2-weighted. However, at these times it is precisely the substances with short transversal relaxation times T2 that have already significantly decayed. The contrast in UEZ sequences is therefore limited to T1 or PD contrast (PD=proton density), wherein the contrast from the steady state (after the transient process) of the measurement is provided and depends on the flip angle and the repetition time. In particular, with the present UEZ sequences it is consequently also not possible to generate T2 contrasts in which substances with long T2 are markedly brighter in the image than substances with short T2, as they are required for many clinical questions.